The field of the invention is magnetic resonance imaging (“MRI”) systems and methods and, more particularly, the invention relates to methods for non-contrast enhanced magnetic resonance angiography (“MRA”).
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the nuclear spins in the tissue attempt to align with this polarizing field, but process about it in random order at their characteristic Larmor frequency. Usually the nuclear spins are comprised of hydrogen atoms, but other NMR active nuclei are occasionally used. A net magnetic moment Mz is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B1; also referred to as the radiofrequency (RF) field) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped” into the x-y plane to produce a net transverse magnetic moment Mt, which is rotating, or spinning, in the x-y plane at the Larmor frequency. The practical value of this phenomenon resides in the signal which is emitted by the excited spins after the excitation field B1 is terminated. There are a wide variety of measurement sequences in which this nuclear magnetic resonance (“NMR”) phenomenon is exploited.
When utilizing these signals to produce images, magnetic field gradients (Gx, Gy, and Gz) are employed. Typically, the region to be imaged experiences a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The emitted MR signals are detected using a receiver coil. The MRI signals are then digitized and processed to reconstruct the image using one of many well-known reconstruction techniques.
Magnetic resonance angiography (MRA) and, related imaging techniques, such as perfusion imaging, use the NMR phenomenon to produce images of the human vasculature or physiological performance related to the human vasculature. There are three main categories of techniques for achieving the desired contrast for the purpose of MR angiography. The first general category is typically referred to as contrast enhanced (CE) MRA. The second general category is phase contrast (PC) MRA. The third general category is time-of-flight (TOF) or tagging-based MRA.
To perform CE MRA, a contrast agent, such as gadolinium, is injected into the patient prior to the magnetic resonance (MR) angiogram to enhance the diagnostic capability of the MR angiogram. Perfusion imaging is employed to assess the viability of tissues. A contrast agent is administered to the subject and a series of MR images are acquired as the contrast agent perfuses into the tissues of interest. From this series of contrast-enhanced MR images hemodynamic parameters such as blood flow, blood volume, and mean transit time may be computed.
While CE MRA is a highly effective means for noninvasively evaluating the vascular and physiological performance, for example, by studying perfusion, the technique suffers from several additional drawbacks. First, the contrast agent that must be administered to enhance the blood vessel carries a significant financial cost. Second, contrast agents such as gadolinium have recently been shown to be causative of a debilitating and potentially fatal disorder called nephrogenic systemic fibrosis (NSF). Third, CE MRA, may not provide accurate or sufficient hemodynamic information, so that it is not always feasible to determine if a stenosis is hemodynamically significant or to assess the perfusion in a clinically useful manner.
Despite strong incentives to move away from CE MRA imaging, current methods for non-contrast angiography are limited in their utility because they are sensitive to patient motion, do not consistently or accurately portray vessel anatomy in patients with severe vascular disease, and require excessively long scan times. For example, while single shot acquisition methods such as two-dimensional (“2D”) balanced steady-state free precession (“bSSFP”) have the potential to reduce motion artifacts and shorten exam times, arterial conspicuity is inadequate due to high background signal. Moreover, bSSFP methods do not lend themselves to the creation of maximum intensity projection (“MIP”) angiograms. In one example, a saturation-recovery bSSFP pulse sequence employed for cardiac perfusion imaging following the administration of a paramagnetic contrast agent is described by W. G. Schreiber, et al., in “Dynamic Contrast-Enhanced Myocardial Perfusion Imaging Using Saturation-Prepared TrueFISP,” JMRI, 2002; 16:641-652. However, this pulse sequence applies a spatially non-selective saturation pulse that suppresses the signal from blood and, thus, cannot be employed for MRA without contrast material. Additionally, Schreiber's method does not provide a means for distinguishing arteries from veins.
Phase contrast (PC) MRA is largely reserved for the measurement of flow velocities and imaging of veins. Phase contrast sequences are the basis of MRA techniques utilizing the change in the phase shifts of the flowing protons in the region of interest to create an image. Spins that are moving along the direction of a magnetic field gradient receive a phase shift proportional to their velocity. Specifically, in a PC MRA pulse sequence, two data sets with a different amounts of flow sensitivity are acquired. This is usually accomplished by applying gradient pairs, which sequentially dephase and then rephase spins during the sequence. The first data set is acquired using a “flow-compensated” pulse sequence or a pulse sequence without sensitivity to flow. The second data set is acquired using a pulse sequence designed to be sensitive to flow. The amount of flow sensitivity is controlled by the strength of the bipolar gradient pairs used in the pulse sequence because stationary tissue undergoes no effective phase change after the application of the two gradients, whereas the different spatial localization of flowing blood is subjected to the variation of the bipolar gradient. Accordingly, moving spins experience a phase shift. The raw data from the two data sets are subtracted to yield images that illustrate the phase change, which is proportional to spatial velocity. To perform PC MRA pulse sequences, a substantial scan time is generally required and the operator must set a velocity-encoding sensitivity, which varies unpredictably depending on a variety of clinical factors.
Fortunately, TOF imaging techniques do not require the use of a contrast agent and do not rely on potentially-precarious velocity encoding sensitivities. Contrary to CE-MRA, which relies on the administered contrast agent to provide an increase in measured MR signal, TOF MRA relies on the inflow of blood into an imaging volume to increase the signal intensity of the vasculature as compared to the stationary background tissues. This is achieved by the application of a number of RF excitation pulses to the imaging volume that cause the magnetization of the stationary background tissues to reach a saturation value. Since inflowing blood entering the imaging volume is not exposed to the same number of RF excitation, it will provide higher MR signal intensity than the background tissue. The differences between the signal intensity of the stationary background tissues and the inflowing blood thus provide a contrast mechanism exploited by TOF MRA.
In an effort to increase contrast attributable to the relatively small signal levels or weight particular signals, for example, those attributable to cerebral blood flow (CBF) or another measurable mechanism, various “tagging” or “labeling” methods have been developed. One such method is referred to as the arterial spin labeling (ASL) family of techniques. These techniques have been used to detect and provide a quantitative measure of neuronal activity.
Two such methods of non-contrast enhanced MRA are described, for example, by M. Katoh, et al., in “Free-Breathing Renal MR Angiography With Steady-State Free-Precession (SSFP) and Slab-Selective Spin Inversion: Initial Results,” Kidney International, 2004; 66:1272-1278, and by Y. Yamashita, et al., in “Selective Visualization of Renal Artery Using SSFP with Time-Spatial Labeling Inversion Pulse: Non-Contrast Enhanced MRA for Patients with Renal Failure,” Proc. Intl. Soc. Mag. Reson. Med. 13 (2005) p. 1715. The method described by Katoh utilizes a three-dimensional (“3D”) acquisition with a pre-inversion of the 3D region, while Yamashita employs two inversion pulses (one spatially selective and the other spatially non-selective). Each of these methods uses inversion preparation pulses rather than saturation pulses and further requires the use of a 3D, rather than 2D, acquisition for MRA. Given the substantial thickness of the 3D imaging slab, inflowing unsaturated spins must travel a large distance (for example, up to several centimeters) to replace in-plane saturated ones. Consequently, there is poor depiction of slowly flowing arterial spins. In fact, the inversion time, TI, must be very long (on the order of 1 second) to provide adequate inflow of even moderately fast flowing arterial spins. The long TI spans both the systolic and diastolic phases of the cardiac cycle. Given the long TI, it is problematic to synchronize data acquisition to diastole. In addition, 3D acquisitions are too time-consuming to permit data acquisition within a single breath-holding period.
Unfortunately, TOF and tagging or labeling methods have additional drawbacks. For example, tagging or labeling methods are generally ill suited for dynamic, time-resolved imaging studies designed to produce a visualization, such as a movie, of the flow propagation within the arteries. For example, the large imaging volumes used in 3D labeling techniques result in the inflowing spins becoming saturated across the imaging volume, diminishing attainable image contrast. As such, visualization of the propagation of flow within arteries has traditionally been accomplished with the use of CE MRA or PC MRA. CE MRA has the above-noted drawbacks. Within the context of dynamic, time-resolved angiographic studies, PC MRA is particularly time consuming, typically, requiring a minute of longer to collect one slice location, which precludes rapid imaging of a large vascular territory. Additionally, phase-contrast imaging requires pre-selection of velocity encoding sensitivity and specialized processing of the phase-information of the MR images; the latter of which is prone to errors stemming from phase aliasing, random phase in regions of low signal intensity, concomitant gradients, and eddy current effects.
Therefore, it would be desirable to have a system and method for performing angiographic studies using MRI systems without the drawbacks presented by CE-MRA, PC-MRA, TOF, or traditional labeling/tagging techniques. Furthermore, it would be desirable to have a system and method for MR angiography that allows the user to image a volume to produce a dynamic or time-resolved series of images of the volume.